The present invention relates to scintillation cameras. In particular, the invention relates to a method and apparatus for improving the quality of images produced during positron emission tomography.
In the human body, increased metabolic activity is associated with an increase in emitted radiation. In the field of nuclear medicine, increased metabolic activity within a patient is detected using a radiation detector such as a scintillation camera.
Scintillation cameras are well known in the art, and are used for medical diagnostics. A patient ingests, inhales or is injected with a small quantity of a radioactive isotope. The radioactive isotope emits gamma rays that are detected by a scintillation medium in the scintillation camera. The scintillation medium is commonly a sodium iodide crystal, BGO or other. The scintillation medium emits a small flash or scintillation of light, in response to stimulating radiation, such as from a patient. The intensity of the scintillation of light is proportional to the energy of the stimulating photon, such as a gamma photon. Note that the relationship between the intensity of the scintillation of light and the gamma ray is not linear.
A conventional scintillation camera such as a gamma camera includes a detector which converts into electrical signals gamma rays emitted from a patient after radioisotope has been administered to the patient. The detector includes a scintillator and photomultiplier tubes. The gamma rays are directed to the scintillator which absorbs the radiation and produces, in response, a very small flash of light. An array of photodetectors, which are placed in optical communication with the scintillation crystal, converts these flashes into electrical signals which are subsequently processed. The processing enables the camera to produce an image of the distribution of the radioisotope within the patient.
Scintillation cameras are used to take four basic types of pictures: spot views, whole body views, partial whole body views, SPECT views, and whole body SPECT views.
A spot view is an image of a part of a patient. The area of the spot view is less than or equal to the size of the field of view of the gamma camera. In order to be able to achieve a full range of spot views, a gamma camera must be positionable at any location relative to a patient.
One type of whole body view is a series of spot views fitted together such that the whole body of the patient may be viewed at one time. Another type of whole body view is a continuous scan of the whole body of the patient. A partial whole body view is simply a whole body view that covers only part of the body of the patient. In order to be able to achieve a whole body view, a gamma camera must be positionable at any location relative to a patient in an automated sequence of views.
The acronym xe2x80x9cSPECTxe2x80x9d stands for single photon emission computerized tomography. A SPECT view is a series of slice-like images of the patient. The slice-like images are often, but not necessarily, transversely oriented with respect to the patient. Each slice-like image is made up of multiple views taken at different angles around the patient, the data from the various views being combined to form the slice-like image. In order to be able to achieve a SPECT view, a scintillation camera must be rotatable around a patient, with the direction of the detector head of the scintillation camera pointing in a series of known and precise directions such that reprojection of the data can be accurately undertaken.
A whole body SPECT view is a series of parallel slice-like transverse images of a patient. Typically, a whole body SPECT view consists of sixty four spaced apart SPECT views. A whole body SPECT view results from the simultaneous generation of whole body and SPECT image data. In order to be able to achieve a whole body SPECT view, a scintillation camera must be rotatable around a patient, with the direction of the detector head of the scintillation camera pointing in a series of known and precise directions such that reprojection of the data can be accurately undertaken.
Therefore, in order that the radiation detector be capable of achieving the above four basic views, the support structure for the radiation detector must be capable of positioning the radiation detector in any position relative to the patient. Furthermore, the support structure must be capable of moving the radiation detector relative to the patient in a controlled manner along any path.
In order to operate a scintillation camera as described above, the patient should be supported horizontally on a patient support or stretcher.
A certain design of gantry or support structure for a scintillation camera includes a frame upon which a vertically oriented annular support rotates. Extending out from the rotating support is an elongate support. The elongate generally comprises a pair of arms. The pair of arms generally extends through a corresponding pair of apertures in the rotating support. One end of the pair of arms supports the detector head on one side of the annular support. The other end of the pair of arms supports a counter balance weight. Thus, the elongate support is counterbalanced with a counterweight on the opposite side of the detector head.
With such a design of support suture for a scintillation camera, a patient must lie on a horizontally oriented patient support. The patient support must be cantilevered so that the detector head can pass underneath the patient. If the detector head must pass underneath only one end of the patient, such as the patent""s head, the cantilevered portion of the patient support is not long enough to cause serious difficulties in the design of the cantilevered patient support. However, if the camera must be able to pass under the entire length of the patient, the entire patient must be supported by the cantilevered portion of the patient support. As the cantilevered portion of the patient support must be thin so as not to interfere with the generation of images by the scintillation camera, serious design difficulties are encountered.
Among the advantages associated with such as design of support structure is that a patient may be partially pass through the orifice defined by the annular support so that the pair of arms need not be as long. However, the patient support must be able to support the patient in this position relative to the annular support, must be accurately positionable relative to the annular support, and must not interfere either with the rotation of the annular support or with the cables which will inevitably extend from the detector head to a nearby computer or other user control.
The photomultiplier tubes in a scintillation camera generate electric signals. The signals are processed, and images are created corresponding to the radiation emitted by the patient.
From time to time, images are generated that contain one or more artifacts or flaws. Artifacts are often caused by one or more malfunctioning photomultiplier tubes. A malfunctioning photomultiplier tube maybe generating incorrect signals, may be generating no signal at all, or the processing of the signals from a particular photomultiplier tube may not be proper.
To determine the cause of the artifact and then correct the artifact, it is important to identify all malfunctioning photomultiplier tubes. However, inspecting and testing photomultiplier tubes is difficult, time consuming and expensive.
From time to time, images of poor quality are also generated. Of particular concern are the images produced by Position Emission Tomography. Position Emission Tomography (PET) is a practice common in the art wherein two detectors are placed with their fields of view at 180xc2x0 to one another. After the patient ingests the isotope, positrons are emitted from areas where is isotope has gathered in the body. The positrons that are released from the body in opposite directions collide with electrons in the body and effectively form two gamma rays.
The gamma rays are detected by the detectors and as mentioned above are used to generate images. However, in PET, only gamma rays originating from a collision between a positron and an electron that are detected at 180xc2x0 (referred to as coincidence) from one another are considered true events. Preferably only true events are used to generate images.
Unfortunately what sometime occurs is that the gamma ray will ricochet off a second electron in the body before being emitted and the angle is changed. The two gamma rays will not be detected at 180xc2x0 from one another, resulting in a xe2x80x9crandomxe2x80x9d event. Random events are really just noise signals that when used to generate an image, cause poor quality imagery. It is known in the art that an increase in area (of field of view) results in an increase in the probability of random events. Since conventional PET cameras use relatively large detectors with large fields of view and they commonly use the total data values for the entire detector head the chance of using random events to generate an image is high. As well, since data from a large field of view must be processed, the time frame window during which data is analysed is large resulting in yet a higher probability of detecting random events.
In Constant Fraction Discrimination (CFDs) cameras, the probability of random events is also relatively high resulting in poorer quality images. FIG. 1 illustrates the data obtained from a Constant Fraction Discriminator. Constant Fraction Discriminators use a constant fraction (or percentage) of the input pulse to precisely determine the timing of an event. Inaccuracies occur when two events are detected in such a short time frame such as to create overlap. In the data when two or more events overlay, it is impossible to separate them to obtain before an event in order to separate the data as seen in FIG. 1, the data from areas A, B and C can be separated in order to analyse the individual events 1 and 2.
An object of the invention is to provide a method and apparatus for improving a PET image quality. This is achieved by analysing individual photomultiplier tubes for true events and by providing time stamps to photomultiplier tube signals. Analysing data from individual photomultiplier tubes as opposed to entire detector field of views results in smaller areas and smaller amounts of data to be processed. This then permits smaller time frame windows to be used. The use of time stamps also permits data before and after a particular event to be kept as record.
In a positron emission tomography (PET) study using a scintillation detector wherein the scintillation detector has a scintillation crystal and a plurality of photomultiplier tubes and a pair of coincident gamma events due to an annihilation of a position are detected by the scintillation detector in order to locate the positron emission, a method for identifying the coincident gamma events is provided according to one aspect of the present invention. The method comprises the steps of: (a) receiving a photomultiplier tube signal from the photomultiplier tube when a gamma event occurs; (b) digitising the photomultiplier tube signal; (c) generating a clock signal providing a time stamp for the photomultiplier tube signal; and (d) generating an encoded signal using the digitised photomultiplier tube signal and the clock signal, the encoded signal comprising an encoded photomultiplier tube signal followed by an encoded time stamp; wherein, in a subsequent event-positioning process, photomultiplier tube signals caused by coincident gamma events are indentified by means of the encoded time stamp and the identified encoded signals are utilized for positioning of the gamma events.
In a positron emission tomography (PET) scanner using a scintillation detector wherein the scintillation detector has a scintillation crystal and a plurality of photomultiplier tubes and a pair of coincident gamma events due to an annihilation of a position are detected by the scintillation detector in order to locate the positron emission, an apparatus for identifying the coincident gamma events is provided according to another aspect of the invention. The apparatus comprises: (a) a photomultiplier tube for generating a photomultiplier tube signal when a gamma event occurs; (b) an analog-to-digital converter for digitising the photomultiplier tube signal;(c) a clock for generating a clock signal providing a time stamp for the photomultiplier tube signal; and (d) means for generating an encoded signal using the digitised photomultiplier tube signal and the clock signal, the encoded signal comprising an encoded photomultiplier tube signal followed by an encoded time stamp; wherein, in a subsequent event-positioning process, photomultiplier tube signals caused by coincident gamma events are identified by means of the encoded time stamp and the identified encoded signals are utilized for positioning the gamma events.
In a positron emission tomography (PET) scanner using a scintillation detector wherein the scintillation detector has a scintillation crystal and a plurality of photomultiplier tubes and a pair of coincident gamma events due to an annihilation of a position are detected by the scintillation detector in order to locate the positron emission, an apparatus for improving the image is provided according to another aspect of the invention. The apparatus comprises: (a) means for generating a photomultiplier tube signal after an event; (b) means for generating a code signal identifying the photomultiplier tube; (c) means for generating a clock signal providing a time stamp for the photomultiplier tube signal; (d) means for generating an encoded signal comprising the photomultiplier tube signal followed by the code signal and the time stamp; (e) means for determining whether the encoded signal has been caused by a true event; and (e) means for calculating the position of the event using the determined encoded signal.
Other advantages, objects and features of the present invention will be readily apparent to those skilled in the art from a review of the following detailed description of preferred embodiments in conjunction with the accompanying drawings and claims.